MRI systems are widely used for medical diagnosis. They provide high image resolution in advanced applications such as cardiac studies and neuro-scanning. The systems are unique in being able to provide detailed images of soft tissue when bone or other tissue may otherwise obstruct the view.
MRI systems have undergone continual advancement since their inception in the 1970s. Partly through development of larger and more powerful magnets, substantial improvements in image quality have been realized. However, use of these magnets tends to increase the size and cost of MRI systems.
As a premier medical imaging and diagnostic tool, cost is a significant impediment to wider deployment of MRI systems throughout the world. Other factors limiting greater commercialization of these systems include requirements for a large, stable power supply, and size and weight.
For many medical applications MRI systems should provide magnetic field strengths ranging from 0.3 Tesla up to at least 3 Tesla, although even larger field strengths (e.g., up to 20 Tesla or more) may be used in research. Magnetic fields of such intensity are generally formed with large superconducting magnetic coils (traditionally Low Temperature Superconductor (LTS) coils). To operate the magnet in a superconducting mode, the Low Temperature Superconductor (LTS) coils must operate in the vicinity of 4.5 K, i.e., a temperature below the critical temperature of the coil material. This generally requires enclosure of the LTS coils within a vessel cooled with liquid helium.
Although superconducting MRI systems are smaller and more efficient than magnets that operate with coils formed of resistive wire, they remain large and bulky in the absolute sense. Thus, it continues to be a desire in the art to reduce the mass and volume of MRI systems while not sacrificing the strength and efficiency of the magnets.
In addition to providing a main, relatively uniform and stable magnetic field across an imaging volume, the MRI system may include a set of low power gradient magnets ranging in field strength from 18 to 27 millitesla. The coils of the gradient magnets are cooperatively pulsed to provide local variations in the field strength. This enables selection of views within portions of the image volume.
When me gradient coils are electrically pulsed, the resulting time changing magnetic flux induces eddy currents within conductive components positioned about the imaging volume. These, in turn, produce secondary magnetic fields that can degrade the quality and effectiveness of the field set up by the gradient coils and the main magnet. To partially compensate for this effect, a set of shield gradient coils may be pulsed to set up fields which counter those portions of the gradient coil field that extend into the helium-cooled vessel. Although the shield gradient coils can effectively cancel a portion of the field generated by the gradient coils, there remains a need to further suppress the resulting eddy currents.
The shield gradient components and other shield components, e.g., thermal radiation shields, add cost and weight to MRI systems by increasing the diameter of the magnet main coils. Yet it is generally desirable to provide lower-cost, easily deployable MRI systems without sacrificing image quality. For MRI systems to have broader application in medical diagnostics, the size, weight and cost associated with operating the system should be further reduced.
The invention provides a Magnetic Resonance Imaging (MRI) system having a vacuum vessel positioned about an imaging volume, one or more high temperature superconducting coils positioned within the vacuum vessel, and a cryocooler coupled to the vacuum vessel to operate the superconducting coil at a temperature above 10 K. At least one gradient coil is positioned between an imaging volume and the superconducting coil without any thermal shielding interposed between the gradient coil and the superconducting coil.
The invention also provides a method of forming an MRI system including forming at least one winding of the main field generating coils with high temperature superconducting material, positioning the winding in a vessel for receiving cryogenic fluid, and positioning a gradient coil between the imaging volume and the winding without placing a thermal radiation shield between the gradient coil and the winding.
Wherever appropriate, like reference numbers are used throughout the figures to refer to like parts.